Protein detector based on molecular controlled semiconductor resistor

ABSTRACT

The invention provides a semiconductor device for the detection of an active site-containing protein or a ligand thereof in a solution, said device comprising at least one insulating or semi-insulating layer; at least one conducting semiconductor layer, two conducting pads on top of the upper layer making electrical contact with said at least one conducting semi-conductor layer, such that electrical current can flow between them at a finite distance from the surface of the device; a protective molecular layer fabricated on top of said upper layer and protecting said layer from corrosion; and said ligand or active site-containing protein linked to said protective molecular layer. Exposure of said ligand or active site-containing protein to a solution containing said active site-containing protein or ligand, respectively, causes a current change through the device when a constant electric potential is applied between the two conducting pads. The semi-conductor device can be seen as a molecularly controlled semiconductor resistor (MOCSER) protein sensor based on doped and undoped GaAs stack structure. The GaAs is protected against etching in aqueous environments by the protective molecular layer.

TECHNICAL FIELD

The present invention relates to semiconductor devices for detectionand/or quantification of proteins, more particularly activesite-containing proteins, or ligands thereof, more specifically to suchdevices based on molecular controlled semiconductor resistors.

BACKGROUND ART

The detection of covalent and noncovalent binding events betweenmolecules and biomembranes is a fundamental goal of contemporarybiochemistry and analytical chemistry. This detection serves for thebasic study of central biological processes like signaling, and for thedevelopment of high throughput screening of drug candidates from largelibraries of molecules that potentially recognize a specific membranereceptor. Currently, such studies are performed routinely usingfluorescence methods (Chattopadhyay and Raghuraman, 2004),surface-plasmon resonance (SPR) spectroscopy (Baciu et al., 2008), andelectrochemical methods (Thompson and Krull, 1982; Thompson et al.,1983; Umezawa et al., 1988; Woodhouse et al., 1999; Xu and Bakker, 2009;Dumas et al., 2011; Coldrick et al., 2011). However, there is still needfor novel sensitive miniaturizable detection methods, e.g., forpoint-of-care testing (POCT).

The preparation and characterization of model membranes on solidsupports, e.g., semiconductors, is a practical and scientificallyimportant research area (Tanaka and Sackmann, 2005). Practicalapplications include smart biosensor devices for studying basic membraneprocesses and membrane-analyte interactions, as well as for otherbiotechnological applications (Bieri et al., 1999; Sackmann and Tanaka,2000; Sapuri et al., 2002; Yip et al., 2002).

Recent advances in microelectronics and nanotechnology, improvement insensor function, and emergence of new types of biosensors have increasedthe interest in development of lipid membrane-based systems.Electrochemical methods were applied since they allow direct conversionof biological information to electronic signal. They are well suited forinvestigation of biomembrane functions due to their operationsimplicity, low cost, and capability of real-time measurements.Typically, electrochemical biosensors employ amperometric,potentiometric, or impedimetric transducers (Thompson and Krull, 1982;Thompson et al., 1983; Umezawa et al., 1988; Woodhouse et al., 1999; Xuand Bakker, 2009; Dumas et al., 2011; Coldrick et al., 2011).

Sensors based on field-effect transistor (FET) configuration have beenutilized since the early 1970s (Bergveld, 1972; Bergveld et al., 1978).This special class of sensors makes use of the potentiometric effect ata gate electrode (Thevenot et al., 2001). Currently, biosensingapplications focus on ion-selective FET (ISFET or CHEMFET) devices. InISFET, the regular gate is placed in a liquid electrolyte, and thediffusion of specific analytes toward the electrode can be controlled byinsertion of a selective membrane positioned on the gate. ISFET approachwas utilized, e.g., for studying enzyme-substrate recognition and fordetecting neurons or living cell activity (Baumann et al., 1999;Kharitonov et al., 2000; Schoning and Poghossian, 2002; Bergveld, 2003;Janata, 2004). A theoretical model for biorecognition of acetylcholineapplying enzyme-modified ISFET was provided recently (Goykhman et al.,2009). According to this model, the electrical response of the device,during enzyme-substrate recognition events, depends on cooperativeeffects of local pH changes and molecular dipole variations.

International Patent Publication No. WO 98/19151 (corresponding to U.S.Pat. No. 6,433,356) of the same applicant of the present invention,herewith incorporated by reference in its entirety as if fully disclosedherein, describes a hybrid organic-inorganic semiconductor device andsensors based thereon, said device characterized by being composed of:(i) at least one layer of a conducting semiconductor; (ii) at least oneinsulating layer, (iii) a multifunctional organic sensing moleculedirectly chemisorbed on one of its surfaces, said multifunctionalorganic sensing molecule having at least one functional group that bindsto the said surface of the electronic device, and at least one otherfunctional group that serves as a sensor, and (iv) two conducting padson the top layer making electrical contact with the electricallyconducting layer, such that electrical current can flow between them ata finite distance from the surface of the device. The semiconductordevices disclosed in WO 98/19151 are referred to as molecular controlledsemiconductor resistors (MOCSERs) and described as light or chemicalsensors.

SUMMARY OF INVENTION

It has been found, in accordance with the present invention, that adevice based on the molecular controlled semiconductor resistor (MOCSER)previously described (Gartsman et al., 1998; Vilan et al., 1998; Wu etal., 2000; Rei Vilar et al., 2006) and disclosed in the aforesaid WO98/19151, when covered with a protective molecular layer fabricated ontop of its upper layer that is either a conducting semiconductor layeror an insulating or semi-insulating layer, protecting said upper layerfrom corrosion, in particular, a protective molecular layer comprisingan alkoxysilane-based polymer, can be used for detection of activesite-containing proteins or ligands thereof in a solution, and can thusbe utilized for monitoring processes occurring on a membrane and theinteraction of an active site-containing protein in solution with aligand thereof linked to said protective molecular layer, or vice versa.

In one aspect, the present invention thus relates to a semiconductordevice for the detection of an active site-containing protein or aligand thereof in a solution, said device being composed of at least oneinsulating or semi-insulating layer, at least one conductingsemiconductor layer, two conducting pads, a protective molecular layer,and said ligand or active site-containing protein,

wherein said at least one conducting semiconductor layer is on top ofone of said insulating or semi-insulating layers, said two conductingpads are on both sides on top of an upper layer which is either one ofsaid conducting semiconductor layers or another of said insulating orsemi-insulating layers, making electrical contact with said at least oneconducting semiconductor layer, said protective molecular layer isfabricated on top of said upper layer protecting said upper layer fromcorrosion, and said ligand or active site-containing protein is linkedeither directly or indirectly to said protective molecular layer,

wherein exposure of said ligand or active site-containing protein, to asolution containing said active site-containing protein or ligand,respectively, causes a current change through the semiconductor devicewhen a constant electric potential is applied between the two conductingpads.

In certain embodiments, the ligand or active site-containing protein isdirectly linked to the protective molecular layer of the semiconductordevice of the present invention. In other embodiments, the ligand oractive site-containing protein is indirectly linked to said protectivemolecular layer via a mono- or bi-layer membrane comprising anamphiphilic compound or a mixture thereof, wherein said mono- orbi-layer membrane is adhered to said protective molecular layer. Infurther embodiments, the ligand or active site-containing protein isindirectly linked to said protective molecular layer via a linker suchas a ligand-binding protein, biotin, or a biotin-like molecule.

The semiconductor device of the present invention may further be usedfor quantification of said active site-containing protein or ligandthereof in the solution, wherein the current change through thesemiconductor device when a constant electric potential is appliedbetween the two conducting pads is proportional to the concentration ofsaid active site-containing protein or ligand thereof in the solution.

In another aspect, the present invention provides a method for detectionof an active site-containing protein or a ligand thereof in a solution,said method comprising:

-   -   (i) exposing a semiconductor device as defined above to said        solution; and    -   (ii) monitoring the presence of said active site-containing        protein or ligand in said solution according to the changes in        the current measured in said semiconductor device when a        constant electric potential is applied between the two        conducting pads.

BRIEF DESCRIPTION OF DRAWINGS

FIG. 1 shows a schematic representation of the GaAs pseudomorphic HighElectron Mobility Transistor (pHEMT) structure used for the MOCSERfabrication.

FIGS. 2A-2B show atomic force microscopy (AFM) images showing the effectof water exposure on a bare-GaAs surface (2A) and on a GaAs surfacecovered by −25 nm MPS layer (2B).

FIG. 3 shows fluorescence images of an MPS-APS modified GaAs surface anda glass slide (etched in HF) after exposure to 50 nm or 100 nm EPC-LRBPE(99:1) vesicle-containing solutions and incubation for 5 min. Indicationfor unruptured vesicles were observed on the MPS-APS-modified GaAssubstrates when 100 nm EPC-LRBPE (99:1) vesicles were used (panel A),while the same vesicles formed a bilayer membrane on a glass slide,manifested as a much more uniform fluorescence distribution (panel B).Using smaller EPC-LRBPE (99:1) vesicles (50 nm) led to a uniform bilayermembrane formation on both MPS-APS-modified GaAs substrates (panel C)and glass slides (panel D). Scale bar: 35 μm.

FIG. 4 shows the effect of vesicle size on vesicle rupture when adsorbedon ˜25-nm MPS-APS-modified GaAs devices, observed by 2D- and 3D-AFMsurface-analysis measurements. 100-nm EPC vesicles (marked by blackarrows) were trapped in ˜100-120 nm holes within the MPS polymer,preventing them from attaching to nearby vesicles and from their rupture(panels A-C). Downsizing the EPC vesicles from 100 nm to 50 mm increasedthe possibility of two vesicles to adsorb to the same hole, leading totheir rupture into a homogeneous bilayer membrane formation (panelsD-F).

FIG. 5 shows a 3D fluorescence image of MPS-APS modified GaAs surface.Using spotting technique, the surface was modified at different regionswith Protein G (Prot. G), BSA, and fluorescent tagged sheep anti-humanhemoglobin antibodies (Ab). The highest Ab adsorption was observed whenthe surface was modified with Protein G and non-binding sites wereblocked with BSA (a). Ab adsorption was low when the surface is bare (d)or modified with Protein G only (c), and almost no Ab adsorption wasobserved when the surface was modified with BSA only (b), demonstratingthe blocking ability of BSA. This image further shows the depth profileindicating the number of molecular layers constructed.

FIG. 6 shows normalized change in the MOCSER source-drain current as afunction of time after sequentially adsorbing HEPES buffer (50 mM),Protein G (0.2 mg/ml), BSA (0.1 mg/ml), and anti-human hemoglobinantibodies (0.1 mg/ml), under 0.02 ml/min flow rate. The green arrowsindicate the time when the device was exposed to the respectivesolutions as indicated in the graph, and the red arrows indicate theexposure of the device to phosphate buffer (50 mM) for washing.

FIGS. 7A-7D show a schematic representation of the experimental setupused in the studies described in Example 1. (7A) A peristaltic pump wasused to transfer analyte samples to a GaAs-based MOCSER on top of whicha polydimethylsiloxane (PDMS)-based flow cell was constructed. AnAg/AgCl reference (Ref.) electrode was connected via a salt bridge. (7B)All electrical measurements were performed with wire bonded devices. Thechip contained 16 devices from which 4 were selected and measuredsimultaneously. (7C) a schematic structure of the MPS and APS layers.(7D) The GaAs-based device (1; S: source; D: drain) was coated with3-mercaptopropyltrimethoxysilane (MPS) layer and3-aminopropyltrimethoxysilane (APS) (2), on top of which a lipid bilayermembrane was formed (3), and interactions with various analytes (4) wereinvestigated.

FIG. 8 shows a schematic representation of the experimental setup usedin the studies described in Example 2. Syringe pump was used to transferanalyte samples to a GaAs-based MOCSER on top of which a PDMS-basedmicrofluidic flow cell was constructed. An Ag/AgCl reference (Ref.)electrode was connected via a salt bridge. Sheep anti-human hemoglobinantibodies (Hb Ab) were attached to the MPS-APS modified GaAs surfacethrough Protein G, followed by BSA blocking of the non-binding sites.Hb—human hemoglobin.

FIGS. 9A-9B show normalized change in the MOCSER source-drain current asa function of time when sequentially exposing the EPC-based membraneadsorbed on the device to a phosphate buffer solution (0.05 M) at a pHranging between pH 6.0 to 8.0 under a laminar flow of 0.02 ml/min (9A);and normalized response curve based on the derivative (slope) of thechange in the signal in the peaks' half-maximum for different pHsolutions (9B).

FIGS. 10A-10B show (i) normalized change in the MOCSER source-draincurrent as a function of time when membrane-coated GaAs devices wereexposed to various concentrations of L-glutamic acid (GLU). A rapidincrease in the current was observed upon injection of GLU, and thecurrent decreased when the GLU was washed out by phosphate buffer (10A);and (ii) a plot of the change in the normalized current (the derivativeof the signal) using either MPS-APS-modified GaAs devices or devicesadditionally covered by an EPC-membrane (10B).

FIGS. 11A-11B show (i) normalized change in the MOCSER source-draincurrent as a function of time when membrane-coated GaAs devices wereexposed to various concentrations of L-lysine. A rapid decrease in thecurrent was observed in the presence of L-lysine, and the currentincreased when the L-lysine was washed out by phosphate buffer (11A);and (ii) a plot of the change in the normalized current (the derivativeof the signal) in the case of MPS-APS-modified GaAs devices and thedevices additionally covered by EPC-membrane.

FIG. 12 shows schematic illustration of the strategy used for surfacemodification with streptavidin (panel A) and rabbit anti-streptavidinantibody (panel B) for detection by the devices coated with abiotinylated membrane (EPC-BCPE (8:2)).

FIGS. 13A-13B show the response of a MOCSER device coated withbiotin-containing EPC-BCPE (8:2) membrane upon exposure to streptavidin(13A) and avidin (13B) molecules. In both cases, the indicated amount ofmaterial was dissolved in 100 μl.

FIGS. 14A-14D show (i) normalized change in the MOCSER source-draincurrent as a function of time when exposed to different concentrationsof rabbit anti-streptavidin antibody (whole antiserum) and rabbit serumalone (no antibodies). After exposure to serum, the signal recovered tothe baseline when washed with a phosphate buffer, while there was anon-reversible baseline offset in case of exposure to theanti-streptavidin antibody solution, resulting from specific interactionbetween immobilized streptavidin and rabbit anti-streptavidin antibodies(14A); (ii) normalized change in current when the device was exposed todifferent concentrations of mouse anti-streptavidin antibody (purified).In this case, the signal exhibited non-reversible baseline offset uponwashing with the buffer solution (14B); (iii) calibration curve forresponse due to strong interaction of rabbit anti-streptavidin antibodywith streptavidin, compared with the weak interaction of the sameantibodies with non-biotinylated membrane (14C); and (iv) calibrationcurve for response due to interaction of mouse anti-streptavidinantibody with streptavidin (14D).

FIG. 15 shows a scheme of the double layer formed on the surface of theMOCSER. The MPS layer is not in scale, and in reality, it is about ahundred times thicker than the bilayer membrane. In the case of highconcentration of cations in the double layer, the MOCSER surface isnegatively charged and the current through the MOCSER decreases, whilefor high anion concentration, the MOCSER surface becomes positivelycharged and the current through the MOCSER increases.

FIG. 16 shows normalized change in the MOCSER source-drain current as afunction of time when sequentially exposed to hemoglobin (Hb) dissolvedin phosphate buffer (50 mM), at the concentrations indicated in thegraph under 0.02 ml/min flow rate. The gradient of response (normalizedresponse) is shown in bold.

FIG. 17 shows normalized change in the MOCSER source-drain current as afunction of time when sequentially exposed to hemoglobin (Hb) dissolvedin urine, at the concentrations indicated in the graph under 0.02 ml/minflow rate. The gradient of the response (normalized response) is shownin bold.

FIG. 18 shows a calibration plot representing the normalized change inthe current measured in the MOCSER upon exposure to differentconcentrations of hemoglobin in phosphate buffer (50 mM) and in urine,based on the gradient of the change in the source-drain current as afunction of time.

FIG. 19 shows normalized change in the MOCSER source-drain current as afunction of time when sequentially exposed to human hemoglobin (Hemo;0.5 mg/ml or 0/25 mg/ml) dissolved in urine and phosphate buffer,respectively, without sheep anti-human hemoglobin antibodies attached tothe MPS-APS modified surface. Hemoglobin was introduced as indicated andthere was no response of the MOCSER when anti-hemoglobin antibodies werenot present.

FIG. 20 shows normalized change in the MOCSER source-drain current as afunction of time when sequentially exposed to hemoglobin (Hemo) andavidin as indicated. While in response to hemoglobin, the MOCSERresponded as shown, no response was shown after exposure to avidin,indicating the specificity of the device.

DETAILED DESCRIPTION OF THE INVENTION

In one aspect, the present invention provides a semiconductor devicebased on a molecular controlled semiconductor resistor (MOCSER) for thedetection of an active site-containing protein or a ligand thereof in asolution, as defined above.

In one particular such aspect, the present invention provides asemiconductor device for the detection of said active site-containingprotein in said solution, wherein said semiconductor device comprisessaid ligand linked either directly or indirectly to said protectivemolecular layer, and exposure of said ligand to a solution containingsaid active site-containing protein causes a current change through thesemiconductor device when a constant electric potential is appliedbetween the two conducting pads.

In another particular such aspect, the present invention provides asemiconductor device for the detection of said ligand in said solution,wherein said semiconductor device comprises said active site-containingprotein linked either directly or indirectly to said protectivemolecular layer, and exposure of said active site-containing protein toa solution containing said ligand causes a current change through thesemiconductor device when a constant electric potential is appliedbetween the two conducting pads.

The term “active site-containing protein”, as used herein, refers to anon-structural protein including, e.g., an antibody, protein antigen,enzyme, protein substrate or inhibitor, receptor, and lectin. The term“ligand”, as used herein with respect to said active site-containingprotein, refers to an ion, molecule, or molecular group that binds tosaid active site-containing protein as defined above to form a largercomplex. Non-limiting examples of active site-containing protein-ligandpairs include an antibody and its antigen, respectively, or vice versa;an enzyme and either a substrate or inhibitor thereof, respectively, ofvice versa; a receptor and either a protein or organic molecule,respectively, or vice versa; and a lectin and a sugar.

In one embodiment, the semiconductor device of the present invention iscomposed of at least one insulating or semi-insulating layer, oneconducting semiconductor layer, two conducting pads, a protectivemolecular layer, and said ligand or active site-containing protein,

wherein said conducting semiconductor layer is on top of one of saidinsulating or semi-insulating layers, said two conducting pads are onboth sides on top of an upper layer which is either said conductingsemiconductor layer or another of said insulating or semi-insulatinglayers, making electrical contact with said conducting semiconductorlayer, said protective molecular layer is fabricated on top of saidupper layer, and said ligand or active site-containing protein is linkedeither directly or indirectly to said protective molecular layer.

The various conducting semiconductor and insulating or semi-insulatinglayers of the semiconductor device of the present invention are definedas in the basic MOCSER disclosed in the aforesaid WO 98/19151.

In certain embodiments, each one of the conducting semiconductor layersin the semiconductor device of the present invention independently is asemiconductor selected from a III-V and a II-VI material, or mixturesthereof, wherein III, V, II and VI denote the Periodic Table elementsIII=Ga, In; V=As, P; II=Cd, Zn; VI=S, Se, Te. In preferred embodiments,each one of the conducting semiconductor layers is doped GaAs or doped(Al,Ga)As.

In certain embodiments, each one of the insulating or semi-insulatinglayers in the semiconductor device of the present inventionindependently is a dielectric material selected from silicon oxide,silicon nitride or an undoped semiconductor selected from a III-V and aII-VI material, or mixtures thereof, wherein III, V, II and VI denotethe Periodic Table elements III=Ga, In; V=As, P; II=Cd, Zn; VI=S, Se,Te. In preferred embodiments, the undoped semiconductor is undoped GaAsor undoped (Al,Ga)As.

The protective molecular layer of the semiconductor device of thepresent invention is aimed at protecting the device from etching, i.e.,corrosion, in aqueous solutions.

In certain embodiments, the protective molecular layer comprises analkoxysilane-based polymer, i.e., a polymer formed by polymerization ofdialkoxysilanes, trialkoxysilanes or tetraalkoxysilanes, preferablytrialkoxysilanes, each one of said alkoxysilanes having a functionalgroup, a biotinylated form thereof, or a mixture of the aforesaid. Inparticular embodiments, the alkoxysilane-based polymer is formed bypolymerization of dialkoxysilanes or trialkoxysilanes of the generalformula (C₁-C₇ alkyl)₂-Si(OR)₂ or (C₁-C₇ alkyl)-Si(OR)₃, respectively,biotinylated forms thereof, or mixtures of the aforesaid, wherein eachof the Rs independently is a (C₁-C₄)alkyl, preferably methyl or ethyl,and the (C₁-C₇)alkyl group of the trialkoxysilane, or one or two of the(C₁-C₇)alkyl groups of the dialkoxysilane, is substituted at a terminalcarbon atom with a functional group such as mercapto, amino, andhydroxyl; and the (C₁-C₇)alkyl group of the trialkoxysilane, or one ortwo of the (C₁-C₇)alkyl groups of the dialkoxysilane is optionallyfurther interrupted with one or more —NH— groups.

The term “alkyl”, as used herein, typically means a straight or branchedhydrocarbon radical, wherein “(C₁-C₇)alkyl” and “(C₁-C₄)alkyl”particularly refer to such radicals having 1-7 or 1-4 carbon atoms,respectively. Non-limiting examples of such alkyls include methyl,ethyl, n-propyl, isopropyl, n-butyl, sec-butyl, isobutyl, tert-butyl,n-pentyl, 2,2-dimethylpropyl, n-hexyl, n-heptyl, and the like. The term“(C₁-C₇)alkylene” refers to a straight or branched divalent hydrocarbonradical having 1-7 carbon atoms and include, e.g., methylene, ethylene,propylene, butylene, 2-methylpropylene, pentylene, 2-methylbutylene,hexylene, 2-methylpentylene, 3-methylpentylene, 2,3-dimethylbutylene,heptylene, and the like.

In certain particular embodiments, the alkoxysilane-based polymer isformed by polymerization of a mercapto-functional alkoxysilane of thegeneral formula HS—(C₁-C₇)alkylene-SiR(OR)₂ orHS—(C₁-C₇)alkylene-Si(OR)₃, preferably HS—(C₁-C₇)alkylene-Si(OR)₃, abiotinylated form thereof, or a mixture of the aforesaid, wherein eachof the Rs independently is a (C₁-C₄)alkyl, preferably methyl or ethyl.Non-limiting examples of mercapto-functional alkoxysilanes includemercaptomethylmethyldiethoxysilane [(C₂H₅O)₂(CH₃)Si—CH₂—SH],mercaptomethyl methyldimethoxysilane [(CH₃O)₂(CH₃)Si—CH₂—SH],3-mercaptopropylmethyl diethoxysilane [(C₂H₅O)₂(CH₃)Si—(CH₂)₃SH],3-mercaptopropylmethyl dimethoxysilane [(CH₃O)₂(CH₃)Si—(CH₂)₃SH],3-mercaptopropyltrimethoxysilane (MPS) [(CH₃O)₃Si—(CH₂)₃SH],3-mercaptopropyltriethoxysilane [(C₂H₅O)₃Si—(CH₂)₃SH], and biotinylatedforms thereof.

In other particular embodiments, the alkoxysilane-based polymer isformed by polymerization of an amino-functional alkoxysilane of thegeneral formula H₂N—(C₁-C₇)alkylene-SiR(OR)₂ orH₂N—(C₁-C₇)alkylene-Si(OR)₃, preferably H₂N—(C₁-C₇)alkylene-Si(OR)₃, abiotinylated form thereof, or a mixture of the aforesaid, wherein eachof the Rs independently is a (C₁-C₄)alkyl, preferably methyl or ethyl,and the C₁-C₇ alkylene is optionally interrupted with one or more —NH—groups. Non-limiting examples of amino-functional alkoxysilanes includeN¹-(3-(trimethoxysilyl)propyl)ethane-1,2-diamine[(CH₃O)₃Si(CH₂)₃NH(CH₂)₂NH₂],N¹-(3-(triethoxysilyl)propyl)ethane-1,2-diamine[(CH₃CH₂O)₃Si(CH₂)₃NH(CH₂)₂NH₂], 3-aminopropyltrimethoxysilane (APS)[(CH₃O)₃Si(CH₂)₃NH₂], 3-aminopropyl triethoxysilane[(CH₃CH₂OO)₃—Si(CH₂)₃NH₂], 4-aminobutyltriethoxysilane[(CH₃CH₂O)₃—Si(CH₂)₄NH₂], 4-aminobutyltrimethoxysilane [(CH₃O)₃Si(CH₂)₄NH₂],N¹-(3-(dimethoxy(methyl)silyl)-2-methylpropyl)ethane-1,2-diamine[(CH₃O)₂(CH₃)Si—CH₂CH(CH₃)CH₂NH(CH₂)₂NH₂],N¹-(3-(diethoxy(methyl)silyl)-2-methylpropyl)ethane-1,2-diamine[(CH₃CH₂O)₂(CH₃)Si—CH₂CH(CH₃)CH₂NH (CH₂)₂NH₂],aminopropylmethyldimethoxysilane [(CH₃O)₂(CH₃)Si—(CH₂)₃NH₂],aminopropylmethyldiethoxysilane [(C₂H₅O)₂(CH₃)Si—(CH₂)₃NH₂], andbiotinylated forms thereof.

In further particular embodiments, the alkoxysilane-based polymer isformed by polymerization of a mixture of a mercapto-functionalalkoxysilane, e.g., a mercapto-functional alkoxysilane of the generalformula HS—(C₁-C₇)alkylene-SiR(OR)₂ or HS—(C₁-C₇)alkylene-Si(OR)₃,preferably HS—(C₁-C₇)alkylene-Si(OR)₃, a biotinylated form thereof, or amixture of the aforesaid, as defined above; and an amino-functionalalkoxysilane, e.g., an amino-functional alkoxysilane of the generalformula H₂N—(C₁-C₇)alkylene-SiR(OR)₂ or H₂N—(C₁-C₇)alkylene-Si(OR)₃,preferably H₂N—(C₁-C₇)alkylene-Si(OR)₃, a biotinylated form thereof, ora mixture of the aforesaid, as defined above. In one specific suchembodiment, the alkoxysilane-based polymer is formed by polymerizationof a mixture of MPS and APS.

In one particular embodiment exemplified in the studies described herein(FIG. 1), the semiconductor device of the present invention is composedof a first insulating layer of undoped GaAlAs which is on top of a firstconducting semiconductor layer of n-doped GaAs, said first conductingsemiconductor layer is on top of a second insulating layer of undopedGaAlAs which is on top of a third insulating layer of undoped InGaAs,said third insulating layer is on top of a fourth insulating layer ofGaAs, wherein on top of said first insulating layer is a secondconducting semiconductor layer of GaAs on top of which is an upperconducting semiconductor layer of GaAs, and said protective layer isfabricated on top of said upper conducting semiconductor layer. In amore particular embodiment, the protective molecular layer of thissemiconductor device comprises a polymer formed following polymerizationof a mixture of MPS and APS.

In certain embodiments, the semiconductor device of the presentinvention comprises at least one insulating or semi-insulating layereach independently as defined above, at least one conductingsemiconductor layer each independently as defined above, two conductingpads, a protective molecular layer as defined above, and said ligand oractive site-containing protein directly linked to said protectivemolecular layer via a functional group of the alkoxysilane forming theprotective molecular layer, e.g., an amino, mercapto, carboxyl orhydroxyl group of said alkoxysilane.

In other embodiments, the semiconductor device of the present inventioncomprises at least one insulating or semi-insulating layer eachindependently as defined above, at least one conducting semiconductorlayer each independently as defined above, two conducting pads, aprotective molecular layer as defined above, and said ligand or activesite-containing protein indirectly linked to said protective molecularlayer.

In certain particular such embodiments, said ligand or activesite-containing protein is indirectly linked to said protectivemolecular layer via a mono- or bi-layer membrane comprising anamphiphilic compound or a mixture thereof, wherein said membrane isadhered to said protective molecular layer. In certain more particularsuch embodiments, said ligand or active site-containing protein isimmobilized on, i.e., adsorbed to, or incorporated into, said mono- orbi-layer membrane, e.g., by linking to particular chemical groups insaid membrane that are capable of forming strong non-covalent orcovalent bonds with said ligand or active site-containing protein.

In other particular such embodiments, said ligand or activesite-containing protein is indirectly linked to said protectivemolecular layer via a linker such as a ligand-binding protein, biotin,or a biotin-like molecule. In certain more particular such embodiments,said ligand or active site-containing protein is indirectly linked tosaid protective molecular layer via a ligand-binding protein. Examplesof ligand-binding proteins include, without being limited to, Protein A,Protein G, avidin, streptavidin, and antibodies. The term “antibodies”,as used herein, refers to polyclonal and monoclonal antibodies of avian,e.g. chicken, and mammals, including humans, and to fragments thereofsuch as F(ab′)₂ fragments of polyclonal antibodies, and Fab fragmentsand single-chain Fv fragments of monoclonal antibodies. The term alsorefers to chimeric, humanized and dual-specific antibodies.

Biotin, also known as Vitamin H or coenzyme R, is a water-solubleB-complex vitamin (vitamin B₇) composed of a ureido(tetrahydroimidizalone) ring fused with a tetrahydrothiophene ring,wherein a valeric acid substituent is attached to one of the carbonatoms of the tetrahydrothiophene ring. The terms “biotin-like molecule”and “biotin-like residue” as used herein refer to any compound or aresidue thereof, respectively, having a biotin-like structure, capableof binding to the tetrameric proteins avidin and streptavidin with adissociation constant (K_(d)) similar to that of biotin, i.e., in theorder of ˜10⁻¹⁵ M. Non-limiting examples of biotin-like molecules arediaminobiotin and desthiobiotin, as well as molecules comprising atetrahydroimidizalone ring fused with a tetrahydrothiophene ring whichis found in biotin, or analogs thereof such as those found indiaminobiotin and desthiobiotin.

The term “biotinylated form”, as used herein with respect to thedialkoxysilanes, trialkoxysilanes or tetraalkoxysilanes forming theprotective molecular layer, or the amphiphilic compounds forming themono- or bi-layer membrane, refers to any of said alkoxysilanes oramphiphilic compounds, respectively, when covalently attached to abiotin residue or to a residue of a biotin-like molecule, e.g., via oneof the functional groups thereof. Biotinylation of alkoxysilanes oramphiphilic compounds as defined above can be conducted using anytechnology or method commonly known in the art.

Protein A is a surface protein originally found in the cell wall ofStaphylococcus aureus, capable of binding immunoglobulins. The proteinis composed of five homologous Ig-binding domains that fold into athree-helix bundle, wherein each domain is capable of binding proteinsfrom many of mammalian species, preferably IgGs. In particular, ProteinA binds the heavy chain with the Fc region of most immunoglobulins andalso within the Fab region in the case of the human VH3 family.

Protein G is an immunoglobulin-binding protein expressed in group C andG Streptococcal bacteria much like Protein A but with differentspecificities. It is a cell surface protein that is commonly used inpurifying antibodies through its binding to the Fc region. Protein G inits natural form also binds albumin; however, because serum albumin is amajor contaminant of antibody sources, the albumin binding site has beenremoved from recombinant forms of Protein G.

Avidin is a homotetrameric biotin-binding protein having four identicalsubunits, produced in the oviducts of birds, reptiles and amphibiansdeposited in the whites of their eggs. Each one of the subunits can bindto biotin with high affinity and specificity, wherein the K_(d) ofavidin is ˜10⁻¹⁵ M, making it one of the strongest known non-covalentbonds.

Streptavidin is a protein purified from Streptomyces avidinii.Streptavidin homo-tetramers have an extraordinarily high affinity forbiotin, wherein its binding to biotin is one of the strongestnon-covalent interactions known in nature.

The ligand or active site-containing protein may be indirectly linked tothe protective molecular layer of the semiconductor device of thepresent invention via a mono- or bi-layer membrane comprising anamphiphilic compound or a mixture thereof, which is adhered to theprotective molecular layer.

In certain embodiments, the amphiphilic compound comprised within saidmonolayer or bilayer membrane is a phospholipid, i.e., a lipid capableof forming a lipid bilayer, a biotinylated form thereof, or a mixture ofthe aforesaid. Such phospholipids may be either phosphoglycerides, alsoknown as glycerophospholipid, or phosphosphingolipids.

Particular types of phosphoglycerides include, without being limited to,plasmalogens; phosphatidates, i.e., phosphatidic acids;phosphatidylethanolamines (cephalin); phosphatidylcholines (lecithin)such as egg phosphatidylcholin (EPC); phosphatidylserine;phospatidylinositol; phosphatidylinositol phosphate, i.e.,phosphatidylinositol 3-phosphate, phosphatidylinositol 4-phosphate, orphosphatidylinositol 5-phosphate, phosphatidylinositol bisphosphate, andphosphatidylinositol triphosphate; glycolipids such asglyceroglycolipids, glycosphingolipids, andglycosylphosphatidylinopsitols; phosphatidyl sugars; and a biotinylatedforms thereof such as dioleoyl-sn-glycero-3-phosphoethanolamine-N-(capbiotinyl) (BCPE), Biotin-Phosphatidylcholine (Cat. No. L-11B16,Echelon®), Biotin Phosphatidylinositol 3-phosphate (Cat. No. C-03B6,Echelon®), Biotin Phosphatidylinositol 4,5-bisphosphate (Cat. No.C-45B6, Echelon®), Biotinylated phosphatidylinositol3,4,5-trisphosphate, and1-((1-octanoyl-N′-biotinoyl-1,6-diaminohexane-2R-octanoyl)phosphatidyl)inositol-3,4,5-triphosphate,tetrasodium salt (PtdIns-(3,4,5)-P₃-biotin (sodium salt); Cayman,Chemical Item Number 10009531).

Examples of phosphosphingolipids include, without being limited to,ceramide phosphorylcholine, ceramide phosphorylethanolamine, ceramidephosphorylglycerol, and biotinylated forms thereof such as BiotinSphingomyelin (Cat. No. S-400B, Echelon®).

The decision whether to link said ligand or active site-containingprotein to said protective molecular layer via a lipid mono- or bi-layermembrane depends on the type and properties of said ligand or activesite-containing protein, wherein formation of such a membrane might bepreferred, e.g., in cases a membrane protein should be linked to theprotective molecular layer as well as in order to avoid non-specificinteractions of either or both of said ligand or active site-containingprotein, and the analyte detected, i.e., said active site-containingprotein or ligand, respectively, with said protective molecular layer.

In certain particular embodiments, the semiconductor device of theinvention comprises at least one insulating or semi-insulating layereach independently as defined above, at least one conductingsemiconductor layer each independently as defined above, two conductingpads, a protective molecular layer as defined above, and said ligand oractive site-containing protein indirectly linked to said protectivemolecular layer via a mono- or bi-layer membrane comprising a mixture ofan amphiphilic compound and a biotinylated form of an amphiphiliccompound, wherein a biotinylated form of said ligand or activesite-containing protein is non-covalently attached via an avidin orstreptavidin molecule to the biotin or biotin-like residues in saidmono- or bi-layer membrane.

In other particular embodiments, the semiconductor device of theinvention comprises at least one insulating or semi-insulating layereach independently as defined above, at least one conductingsemiconductor layer each independently as defined above, two conductingpads, a protective molecular layer as defined above, and said ligand oractive site-containing protein indirectly linked to said protectivemolecular layer via biotin or a biotin-like molecule, wherein saidbiotin or biotin-like molecule is covalently linked to a functionalgroup in said protective molecular layer, and a biotinylated form ofsaid ligand or active site-containing protein is non-covalently attachedvia an avidin or streptavidin molecule to the biotin or biotin-likeresidues linked to said protective molecular layer.

In further particular embodiments, the semiconductor device of theinvention comprises at least one insulating or semi-insulating layereach independently as defined above, at least one conductingsemiconductor layer each independently as defined above, two conductingpads, a protective molecular layer as defined above, and said ligand oractive site-containing protein indirectly linked to said protectivemolecular layer via a ligand binding protein such as Protein A, ProteinG, streptavidin, avidin or an antibody, wherein said ligand bindingprotein is covalently linked to a functional group in said protectivemolecular layer, and non-covalently attached to said ligand or activesite-containing protein.

Ligand binding proteins such as Protein A and Protein G can be used,e.g., when the active site-containing protein indirectly linked to theprotective molecular layer is an antibody. Ligand binding proteins suchas streptavidin and avidin can be used, e.g., to bind biotin or abiotin-like molecule, to which said ligand or active site-containingprotein is linked. An antibody can be used as a ligand binding protein,e.g., when the active site-containing protein indirectly linked to theprotective molecular layer is an antigen capable of forming stronginteractions with said antibody.

The semiconductor device of the present invention may be used for thedetection of an active site-containing protein or a ligand thereof in asolution. Said solution may be an aqueous solution, e.g., aphysiological solution, a bodily fluid such as amniotic fluid, aqueoushumour, vitreous humour, bile, blood serum, breast milk, cerebrospinalfluid, cerumen (earwax), endolymph, perilymph, female ejaculate, gastricjuice, mucus, peritoneal fluid, saliva, sebum (skin oil), semen, sweat,tears, vaginal secretion, vomit and urine, or a bodily fluid-basedsolution, i.e., an aqueous solution in which a bodily fluid isdissolved.

In certain embodiments, the semiconductor device of the presentinvention, in any of the configurations defined above, is further usedfor quantification of said active site-containing protein or ligandthereof in said solution, wherein the current change through thesemiconductor device when a constant electric potential is appliedbetween the two conducting pads is proportional to the concentration ofsaid active site-containing protein or ligand thereof in the solution.

GaAs/AlGaAs-based MOCSER devices as defined above were fabricated asdescribed in detail in the Experimental section hereinafter. Aprotective molecular layer of MPS was fabricated on top of each one ofthe devices, and the thickness and quality of said layer, as well as theprotection against etching provided thereby, were determined. Incontrast to silicon oxide-coated GaAs-based MOCSERs, wherein a dramaticreduction in the sensitivity of the device was observed due to theelimination of the ability to modify the surface states on the GaAs, noreduction in the sensitivity of the MPS-protected GaAs MOCSERs wasobserved.

In one of the studies conducted, an active site-containing protein, moreparticularly, streptavidin or avidin, was bound to the protectivemolecular layer via a lipid bilayer membrane. In order to ensureadhesion of said bilayer membrane to the MPS-coated MOCSER, APS wasfirst adsorbed on top of the MPS-coated GaAs devices.

The bilayer membrane was formed on the MPS-APS coated (modified) GaAsdevices by the vesicle fusion method, using vesicles prepared from eggphosphatidylcholin (EPC); mixtures of EPC and1,2-dioleoyl-sn-glycero-3-phosphoethanolamine-N-(lissamine rhodamine Bsulfonyl) (LRBPE); and mixtures of EPC and1,2-dioleoyl-sn-glycero-3-phosphoethanolamine-N-(cap biotinyl) (BCPE).Vesicles were prepared according to a known protocol, and were thensonicated and down-sized to either 100 or 50 nm, resulting in finalvesicle-concentration of 3×10⁸/μl. For membrane deposition, the MPS-APScoated device was inserted inside a flow system fixed on top of thesensing area of the device, and the mixtures of vesicles were injectedinto the flow cell either manually or using a peristaltic pump. Afterinjection, the vesicles were incubated so as to allow fusion andspreading on the surface to form a lipid bilayer, unfused lipid excesswas removed, and the smoothness and integrity of the bilayer membraneformed were observed.

The effect of vesicle size on the formation of a homogeneous bilayermembrane on the MPS-APS-modified GaAs substrate was evaluated usingfluorescence imaging. As found, incubation of the EPC-LRBPE (99:1)vesicles on the MPS-APS-coated GaAs substrate, for 5 min, was sufficientfor vesicle adhesion and rupture into a bilayer membrane; however, somevesicles remained unruptured, leading to a grainy fluorescence image.Longer incubation time increased the number of unruptured vesicles.Downsizing the EPC-LRBPE (99:1) vesicles from 100 to 50 nm dramaticallydecreased the number of unruptured vesicles on the MPS-APS-coated GaAssubstrate, leading to a homogeneous image similar to that observed onhydrogen fluoride (HF)-etched glass slide. The membrane attachment tothe MPS-coated GaAs devices in the presence of APS was improved comparedto that in the absence of APS. The effect of vesicle size on theformation of a homogeneous bilayer membrane on the MPS-APS-coated MOCSERsubstrate was also evaluated using AFM-surface-analysis measurement. Theroot mean square roughness value of the ˜25 nm MPS-APS-coated GaAsdevices was found to be ˜1.6 nm. The MPS polymer was not homogeneous andcontained some holes varying from 5 nm to 120 nm in diameter, sufficientto trap a single 100 nm vesicle, preventing its attachment to nearbyvesicles and therefore its rupture owing to lack of certain surfacedensity of vesicles. By downsizing the vesicle size from 100 to 50 nm,the possibility of two vesicles to adsorb to the same hole wasincreased, leading to their rupture into a homogeneous bilayer membrane.

The stability and integrity of the bilayer membrane formed over timewere compared to those of a similar membrane formed on HF-etched glass.As found, while integrity of the membrane adsorbed on glass slide wasstable for more than 7 days in the presence of 2 mM CaCl₂, integrity ofthe membrane adsorbed on the MPS-APS-modified GaAs substratedeteriorated after ˜5 days.

After forming a membrane of 50 nm EPC-BCPE vesicles on theMPS-APS-coated MOCSER devices, a solution containing either streptavidinor avidin was added to the biotinylated membrane, allowed to interactwith the biotin molecules in the BCPE, and was then washed; andstreptavidin or avidin attachment to the EPC-BCPE membrane wasevaluated.

In another study conducted, an active site-containing protein, moreparticularly, sheep anti-human hemoglobin antibody, was bound to theprotective molecular layer by immobilizing said antibody on theprotective molecular layer surface via Protein G, acting as a ligandbinding protein, and blocking the non-binding sites by bovine serumalbumin (BSA). In order to enable binding biological molecules to theprotective molecular layer, APS was adsorbed on top of the MPS-coatedGaAs devices.

The change in the current of the MOCSER after sequentially adsorbingProtein G, BSA and anti-human hemoglobin antibodies was tested, and asfound, the net change in current was negative when Protein G wasintroduced into the sensing area; positive upon introducing of BSA; andnegative during anti-human hemoglobin antibodies interaction with thesurface of the device. Since Protein G, BSA, and sheep anti-humanhemoglobin antibodies are all negatively charged protein molecules at pH7.4, these results demonstrate that the sensing mechanism of the MOCSERis different from most generally accepted capacitive theory applicablefor ISFETs.

Example 1 hereinafter describes a study in which the response of abiotinylated lipid bilayer membrane (an EPC-BCPE membrane containing afraction of ˜20% biotin)-coated MPS-APS modified GaAs device tophosphate buffer solutions containing various analytes was tested.Analytes dissolved in phosphate buffer solution were injectedsequentially into a flow cell fixed on top of the sensing area of thedevice, and phosphate buffer solution was injected between the analytesfor washing and removing analyte excess. A constant potential of 1.0 Vwas applied between source and drain of the device, and changes insource-drain current were monitored as function of time. An Ag/AgClpseudo reference electrode was placed in a sealed tube and connected viaa salt bridge to maintain a stable and constant potential over thesurface of the device.

As shown in this Example, whereas the source-drain current response topH change was immediate, stable, and linear within the pH range studied,the changes in the current observed in the bilayer membrane-coateddevice when exposed to various concentrations of negatively- orpositively-charged amino acids at pH=7, exemplified by L-glutamic acidor L-lysine, respectively, were correlated with the analyteconcentration. In particular, while the current increased as theconcentration of L-glutamic acid increased, it decreased with increasingL-lysine concentration. The detection thresholds for L-lysine andL-glutamic acid in the presence of EPC membrane were about 12.5 mM and6.2 mM, respectively, and they improved to 3.2 mM and 1.6 mM forL-lysine and L-glutamic acid, respectively, in the absence of thebilayer membrane, indicating that said membrane reduces the sensitivityof the device by about a factor of four.

Exposing the biotinylated membrane to either streptavidin that isnegatively charged at neutral pH, or avidin that is positively chargedat neutral pH, at concentrations above 0.8 μM at pH=7, resulted in asignificant change in the MOCSER source-drain current. In particular,the current increased as the streptavidin concentration increased, butremained constant when the solution was changed to buffer with nostreptavidin, indicating a strong (and seemingly irreversible on thetime scale of the experiment) binding of the streptavidin to the biotin.When exposed to avidin, the current through the MOCSER was reduced. Whenan EPC-based membrane without biotin was exposed to the same solution, achange in the current was observed; however, this change could becompletely reversed by washing with buffer.

Changes in the source-drain current were observed when devices to whichstreptavidin was initially attached were exposed to rabbitanti-streptavidin antibody in serum. As shown, the current decreasedupon exposure to the antibodies at concentrations of 0.031, 0.125 and 1mg/ml, indicating strong and seemingly irreversible binding of theanti-streptavidin molecules to the biotin-streptavidin complexes,wherein the signal is accumulating as a function of the amount ofanalyte to which the sensor is exposed. As shown in a controlexperiment, when devices comprising an EPC-based membrane without biotinwere exposed to the same antibody containing-serum solution, smallpositive offset in the current was observed upon washing, indicatingnon-specific binding of serum species to the membrane.

The study described in Example 1 demonstrates the sensing of variousspecies that interact differently with the bilayer membrane-coatedMOCSER. In the case of pH and amino acids sensing, the interaction ofthe analytes with the substrate is weak, as validated by the ability toremove the analyte from the sensor by washing with the buffer solution.The dependence of the signal on the analyte concentration in these casesis complicated since it reflects the change in the bilayer charge andthe extent that this change affects the charge on the MOCSER surfaceitself. Importantly, since the molecules forming the membrane arezwitterions, no preference, in terms of the interaction strength, wasobserved in the two oppositely charged amino acids.

Much of the efforts in developing new diagnostic tools are shifting fromdisease diagnostics to disease management (Coughlin et al., 2006). Thismeans that point of care (POC) sensors will play an important role inhelping controlling the patient condition and in evaluating specificmedical treatment. These sensors should be easy to handle and costeffective (Price, 2001; Pflfflin and Schleicher, 2008; Makowski andIvanisevic, 2011). Diseases associated with hemoglobin like anemia,diabetes (Mayer and Freedman, 1983), hematemesis (Ian et al., 2008),hematuria (Landefeld and Beyth, 1993), and hemoglobiuria (Rother, 2005)need continuous monitoring of the hemoglobin over prolong period oftime. Most of the present available POC devices for sensing hemoglobinuse traducers based on amperometric, colorimetric and piezoelectrictechniques (McMurdy et al., 2008; Park et al., 2005) and the mostpopular are the electrode-based sensors which operate on amperometricpotential principles. The major drawback of this technique is that thedevices lose the sensitivity due to over potential applied (Salimi etal., 2005).

Hematuria and hemoglobiuria are diseases associated with hemoglobin andneeds continue monitoring of hemoglobin in urine as they are symptomsfor kidney stones or renal cancer etc (Mayer and Freedman, 1983; Ian etal., 2008; Rother, 2005). Common clinical practices for sensinghemoglobin are by ELISA or dipstick techniques. ELISA providesquantitative results; however, the response time is very long (on hoursscale), it requires expert technician, and specialized equipment whichis both bulky and relatively expensive (Lazcka et al., 2007). The dipstick sensor is simple to use and inexpensive; however, it providesqualitative information only (Messing, 2007).

Example 2 describes a study in which the response of the MPS-APSmodified GaAs MOCSER, wherein sheep anti-human hemoglobin antibodieswere linked to the protective molecular layer via Protein G, to humanhemoglobin dissolved in either phosphate buffer solution or urine wastested. The experimental setup in this study was similar to thatdescribed above. Analytes dissolved in either phosphate buffer solutionor urine were injected sequentially into the flow cell, and eitherphosphate buffer solution or urine was injected between the analytes forwashing and removing analyte excess.

As shown in this Example, the response of the MPS-APS modified GaAsdevice, in which sheep anti-human hemoglobin antibodies were linked tothe protective molecular layer via Protein G, to human hemoglobindissolved in either phosphate buffer solution or urine, was immediateand stable. The current measured in the MOCSER decreased when hemoglobininteracted with the anti-hemoglobin antibodies linked to theMPS-APS-modified surface, and recovered after washing with the phosphatebuffer or urine, wherein the signal was correlated with theconcentration of the analyte molecules. The sensitivity of the device tohemoglobin based on the experiments conducted was 10 μg/ml and 100 μg/mlof hemoglobin in phosphate buffer and urine, respectively. Thesensitivity of the device to hemoglobin in urine was lower than inphosphate buffer, probably as the urine salt concentration is muchhigher.

In order to verify the selectivity of the MOCSER, the source-draincurrent in response to hemoglobin solutions in both phosphate buffer andurine, when sheep anti-human hemoglobin antibodies were not immobilizedon the gate area, and the surface was only functionalized with protein Gand BSA, was measured. As found, no response to the hemoglobin analyteswas observed, indicating the high selectivity of the device. Whenanalytes containing avidin, representing a non-specific antigen, wereintroduced to a MOCSER having a gate area on which sheep anti-humanhemoglobin antibodies were immobilized, no change in the current wasobserved demonstrating the high specificity of the device.

The operation of the MOCSER as a sensor is based on the fact that it iscapacitance sensitive (Ghafar-Zadeh et al., 2010). Thus, when the deviceis immersed in electrolyte solution with a reference electrode, a doublelayer is formed on its surface. Clearly, when the analyte on the surfaceof the membrane is negatively charged, the charge accumulating on thesurface of the device is positive and vice versa. Since the device isbased on n-doped GaAs, positive charge on the surface increases thecharge carrier concentration in the conductive channel and thesource-drain current increases. The opposite is true for negative chargeon the surface of the GaAs that causes depletion in the charge carrierconcentration and hence reduction in the source-drain current.

The amount of charge accumulating on the surface of the GaAs per chargeon the analyte depends on the potential, V, built between the surface ofthe membrane and the GaAs surface which is given by: V∞Qd/∈, where Q isthe charge per surface area of the analyte, d is the thickness of theMPS-APS-membrane layer, and ∈≈2 is the electric permittivity of thislayer. Since most of the thickness is due to the MPS layer, its value isabout the same for a system with or without the membrane. Hence, thedifference between the response curves obtained with and without amembrane stems from increase of ∈ upon coating the device with themembrane. This increase can be due to a thin water layer located betweenthe membrane and the APS layer.

The MOCSER is different from the well-developed ion selective-FET(ISFET) (Mckinley et al., 1984; Fogt et al., 1985), where the gate onthe transistor is replaced with ion selective membrane that allowsspecific ions to penetrate, and these ions define the electric potentialon the gate. It is also different from the regular chemical field-effecttransistor (ChemFET) or insulated gate field-effect transistor (IGFET)where the metal gate terminal is coated with molecules that interactwith a specific analyte (Lee et al., 2009). In these cases, thedetection is performed by monitoring the change in the gate potentialrequired for maintaining a constant current.

In the present device, the current through the device for a givensource-drain potential is determined by the resistivity which iscontrolled by the band bending in the semiconductor: The more bent thebands, the lower the current. In general, the band bending is determinedby the charge on the surface states. Hence, the sensitivity of thedevice stems from the change induced in the surface states charge. Thedensity of surface states of GaAs is of ˜10³ states/cm², and it has beenshown that a change of about 1% in this charge is enough to affect thecurrent in a detectable way (Capua et al., 2009a). The Dipole momentthat would result from this charge reorganization would have very littleeffect on the surface potential and would not be felt by the MOCSER(Naaman, 2011). Hence, the surface states are the source of sensitivityof the current sensor, a sensitivity exceeding what can be obtained bymodifying the voltage on a metal gate.

The approach presented herein, in which chemical interaction occurringon the outer side of the membrane, if present, or the protectivemolecular layer is transformed into an electric signal changing thecurrent measured in the MOCSER can be implemented for studyingprotein-membrane and protein-protein interactions and kinetics.

In summary, the transmission of information through the lipid mono- orbi-layer membrane, if present, or through the protective molecularlayer, on the semiconductor device of the present invention does notinvolve charge/mass transfer, but rather relies on the ability ofadsorbed organic molecules to control electronic devices. Moreparticularly, the operation of this semiconductor is based on variationin the electrochemical potential on the surface of a GaAs FET-likestructure, in which the gate area is covered by chemically adsorbedmolecules. As shown herein, the special sensitivity of the MOCSER stemsfrom the ability to modify the properties of the surface states locatedon the surface of the device. When the adsorbed molecules interact withother species, the energetics of the surface state is modified,affecting the electrochemical potential on the surface and as a resultthe source to drain current of the device.

In another aspect, the present invention thus provides a method fordetection of an active site-containing protein or a ligand thereof in asolution, e.g., an aqueous solution such as a physiological solution, abodily fluid, or a bodily fluid-based solution, said method comprising:

-   -   (i) exposing a semiconductor device as defined above to said        solution; and    -   (ii) monitoring the presence of said active site-containing        protein or ligand in said solution according to the changes in        the current measured in said semiconductor device when a        constant electric potential is applied between the two        conducting pads.

In one particular such aspect, the present invention provides a methodfor detection of said active site-containing protein in said solution,wherein the semiconductor device exposed to said solution in step (i)comprises said ligand linked either directly or indirectly to saidprotective molecular layer, and exposure of said ligand to a solutioncontaining said active site-containing protein causes a current changethrough the semiconductor device when a constant electric potential isapplied between the two conducting pads.

In another particular such aspect, the present invention provides amethod for detection of said ligand in said solution, wherein thesemiconductor device exposed to said solution in step (i) comprises saidactive site-containing protein linked either directly or indirectly tosaid protective molecular layer, and exposure of said activesite-containing protein to a solution containing said ligand causes acurrent change through the semiconductor device when a constant electricpotential is applied between the two conducting pads.

In view of the capabilities of the semiconductor device of theinvention, in certain embodiments, the method of the present inventionmay further be used for quantification of said active site-containingprotein or ligand thereof in the solution, wherein the current changethrough the semiconductor device when a constant electric potential isapplied between the two conducting pads is proportional to theconcentration of said active site-containing protein or ligand thereofin the solution.

The method of the invention may be used, e.g., for studyingreceptor-ligand pair interactions, more particularly, for monitoring theinteraction of a receptor in a solution with a ligand directly orindirectly linked, as defined above, to the protective molecular layer,or vice versa.

In certain embodiments, the active site-containing protein and ligandthereof according to the method of the invention are an antibody and anantigen, e.g., a protein antigen, respectively, or vice versa. Incertain particular embodiments, the semiconductor device used accordingto this method comprises said antibody linked either directly orindirectly to said protective molecular layer, and the method is usedfor selective detection and optionally quantification of said antigen ina solution. In other particular embodiments, the semiconductor deviceused comprises said antigen linked either directly or indirectly to saidprotective molecular layer, and the method is used for selectivedetection and optionally quantification of said antibody in a solution.According to the method of the invention, exposure of said antibody orantigen to a solution containing said antigen or antibody, respectively,causes a current change through the semiconductor device when a constantelectric potential is applied between the two conducting pads. Inparticular such embodiments, the current change through thesemiconductor device is proportional to the concentration of saidantigen or antibody in the solution.

In other embodiments, the active site-containing protein and ligandthereof according to the method of the invention are an enzyme andeither a substrate or inhibitor thereof, respectively, or vice versa. Incertain particular embodiments, the semiconductor device used accordingto this method comprises said enzyme linked either directly orindirectly to said protective molecular layer, and the method is usedfor selective detection and optionally quantification of said substrateor inhibitor in a solution. In other particular embodiments, thesemiconductor device used comprises said substrate or inhibitor linkedeither directly or indirectly to said protective molecular layer, andthe method is used for selective detection and optionally quantificationof said enzyme in a solution. According to the method of the invention,exposure of said enzyme or either substrate or inhibitor to a solutioncontaining either said substrate or inhibitor or said enzyme,respectively, causes a current change through the semiconductor devicewhen a constant electric potential is applied between the two conductingpads. In particular such embodiments, the current change through thesemiconductor device is proportional to the concentration of saidsubstrate or inhibitor or said enzyme in the solution.

In further embodiments, the active site-containing protein and ligandthereof according to the method of the invention are a receptor andeither a protein or organic molecule, respectively, or vice versa. Incertain particular embodiments, the semiconductor device used accordingto this method comprises said receptor linked either directly orindirectly to said protective molecular layer, and the method is usedfor selective detection and optionally quantification of said protein ororganic molecule in a solution. In other particular embodiments, thesemiconductor device used comprises said protein or organic moleculelinked either directly or indirectly to said protective molecular layer,and the method is used for selective detection and optionallyquantification of said receptor in a solution. According to the methodof the invention, exposure of said receptor or either protein or organicmolecule, to a solution containing either said protein or organicmolecule or said receptor, respectively, causes a current change throughthe semiconductor device when a constant electric potential is appliedbetween the two conducting pads. In particular such embodiments, thecurrent change through the semiconductor device is proportional to theconcentration of said protein or organic molecule or said receptor inthe solution.

In still further embodiments, the active site-containing protein andligand thereof according to the method of the invention are a lectin anda sugar, respectively. In certain particular embodiments, thesemiconductor device used according to this method comprises saidlectin, and the method is used for selective detection and optionallyquantification of said sugar in a solution. In other particularembodiments, the semiconductor device used comprises said sugar, and themethod is used for selective detection and optionally quantification ofsaid lectin in a solution. According to the method of the invention,exposure of said lectin or sugar to a solution containing said sugar orlectin, respectively, causes a current change through the semiconductordevice when a constant electric potential is applied between the twoconducting pads. In particular such embodiments, the current changethrough the semiconductor device is proportional to the concentration ofsaid sugar or lectin in the solution.

In certain particular embodiments, the method of the invention is usedfor detection and optionally quantification of hemoglobin or adegradation product thereof in a bodily fluid such as urine or in abodily fluid-based solution, and the semiconductor device exposed tosaid bodily fluid or bodily fluid-based solution comprises an antibodyto said hemoglobin or degradation product thereof linked either directlyor indirectly to the protective molecular layer.

In other particular embodiments, the method of the invention is used fordetection and optionally quantification of a specific protein in abodily fluid or a bodily fluid-based solution, e.g., blood, plasma,urine, saliva, and gastro related solutions, and the semiconductordevice exposed to said bodily fluid or bodily fluid-based solutioncomprises an antibody to said specific protein linked either directly orindirectly to said protective molecular layer.

The invention will now be illustrated by the following non-limitingExamples.

EXAMPLES Experimental Molecular Controlled Semiconductor Resistor(MOCSER) Fabrication

GaAs/AlGaAs MOCSER devices with a 600 μm long and 200 μm wide conductingchannel were fabricated in a standard clean-room by photolithographytechniques based on a pseudomorphic High Electron Mobility Transistor(pHEMT) structure (FIG. 1). Each fabricated chip contained 16 channelsseparated by 200 μm to minimize cross talking and leakage current. Allthe channels were electrically characterized before the measurements,and 4 out of the 16 channels were selected and measured simultaneously.

GaAs Corrosion Protection Using 3-Mercaptopropyltrimethoxysilane (MPS)

Since GaAs is used in an aqueous environment where it is susceptible toetching, a protective layer of MPS was fabricated on top of each deviceaccording to a common procedure (Kirchner et al., 2002). GaAs andGaAs/AlGaAs samples were cleaned in isopropanol, acetone, and ethanolfor 10 min each, followed by UV/ozone oxidation for 10 min. In order toremove oxide layer and to expose the arsenic rich surface, thesubstrates were etched in 2% HF for 5 sec, washed in deionized water(DDW), etched in NH₄OH (25% NH₃) for 30 sec, and washed in DDW again.Immediately after etching, the substrates were dried in nitrogen andimmersed in 15 mM solution of MPS in ethanol at 50° C. for 4 hr.Polymerization of MPS was initiated by adding 3% (volume) of NH₄OH(25%), after which the solution was kept at 50° C. for additional 16 hr.The samples were then rinsed with ethanol and dried under a stream ofnitrogen. The thickness of the MPS layer was verified by ellipsometry(J. A. Woollam, model M-2000V) measurements within a range of 399-1000nm, and estimated to be consistently in a range of 25-30 nm.

Corrosion protection was evaluated by means of atomic force microscopy(AFM) surface-analysis measurements (BioScope AFM, Veeco Metrology LLC,Santa Barbara, Calif.) on GaAs samples with and without MPS layerimmersed in water for 24 hr. AFM images were acquired in tapping mode inwater at room temperature (23-25° C.) using BioScope AFM with NanoscopeIV controller equipped with a large (G) scanner. As shown in FIG. 2, inthe absence of MPS, the water molecules dissolved the oxides formed atthe bare GaAs surface, resulting in continuous etching of the GaAssubstrate; and the ensuing surface roughness was of the order of tens ofμm (FIG. 2A). However, in case of GaAs substrates protected by MPS, nosignificant etching was observed and the roughness was the same as thatof the GaAs wafer before etching, on the order of nanometers (FIG. 2B).

Images of 3×3 μm scan size were recorded using oxide-sharpenedmicrofabricated Si₃N₄ cantilevers (DNP-S, Veeco Metrology, SantaBarbara, Calif.) with a nominal spring constant of ˜0.12 N/m (asspecified by the manufacturer) at scan rates of 1-3 Hz. Three types ofimages were acquired simultaneously: topography, amplitude and phase.The typical target amplitude was 300 mV (˜22 nm) and set point was230-240 mV (˜17 nm) for all measurements. Image analysis was performedusing WSxM 5.0 Develop 1.2 software (Schoning and Poghossian, 2002).

In contrast to silicon oxide-coated GaAs-based MOCSERs, wherein adramatic reduction in the sensitivity of the device was observed due tothe elimination of the ability to modify the surface states on the GaAs,no reduction was observed in the sensitivity of the MPS-protectedGaAs-based MOCSERs.

In a further experiment, MPS solutions of 3 μl MPS in 1 ml EtOH or 4 μlMPS in 1 ml EtOH were used. First layer adsorbed for 4 hours, and afteradding NH₄OH, polymerization starts and continued for 16 hours. AFMimages taken support the conclusion that the quality of the primarylayer depends on both the concentration of the adsorption solution andthe time of deposition. Increasing deposition time from 4 to 8 hourssignificantly reduced the roughness of the surface, while furtherincreasing the deposition time did not add to the surface smoothness.MPS solutions with concentrations of 3 or 4 μl MPS in 1 ml EtOH resultin a continuous thin polymer layer with no pinholes.

Membrane Formation on MPS-Coated GaAs Devices

The bilayer membrane was formed on the MPS-coated GaAs devices by thevesicle fusion method (Richter et al., 2006; Wong et al., 1999;Sackmann, 1996), using the following lipids: egg phosphatidylcholin(EPC); 1,2-dioleoyl-sn-glycero-3-phosphoethanolamine-N-(lissaminerhodamine B sulfonyl) (LRBPE); and1,2-dioleoyl-sn-glycero-3-phosphoethanolamine-N-(cap biotinyl) (BCPE).

Vesicles were prepared according to a protocol described by Barenholz etal. (1977) and Boukobza et al. (2001). Individual EPC lipids, mixturesof EPC-LRBPE (99:1, molar ratio), or mixtures of EPC-BCPE (8:2, molarratio) were first diluted in tert-butyl alcohol and then lyophilized toremove all traces of organic solvent. The dry phospholipids wererehydrated by phosphate buffer containing 2 mM CaCl₂ (5 mg lipid in 450μl of buffer 0.05 M, pH 7.0), creating multilamellar vesicles. In orderto form unilamellar vesicles, the mixture was sonicated for 10 min usinga Heat System Sonifier. Following sonication, vesicles were down-sizedto 100 nm by extruding repetitively (39 times) through a polycarbonatefilm with 100 nm pores (Anatop™, Whatman) or to 50 nm by additionalextrusion (39 times) through a polycarbonate film with 25 nm pores(Anatop™, Whatman). Vesicle size-distribution was measured using dynamiclight scattering size-measurement (Viscotek 802 DLS, MalvernInstruments, Worcestershire, UK) and was found to be about 10% in atypical preparation. Final vesicle-concentration was 3×10⁸/μl. Althoughvesicles were found to be stable for several days in phosphate buffer(0.05 M, pH 7.0), all vesicles were used immediately after preparation.

The characterization of the membrane formation was performed on GaAssubstrate and not on the device itself. In order to ensure the adhesionof the bilayer membrane to the MPS-coated MOCSER, APS was adsorbed ontop of the MPS-coated GaAs and GaAs/AlGaAs samples by overnightevaporation inside a sealed Petri dish at room temperature. Next, a flowcell was constructed on the surface of the MPS-APS-modified GaAs samplesby attaching the HF-etched glass slide to the GaAs substrate withdouble-sided tape. For comparison, a bilayer membrane was formed on anHF-etched glass slide as well, on which a flow cell made of twoHF-etched glass slides was constructed in a similar way. Solutions of 50nm or 100 nm EPC-vesicles, mixtures of 50 nm or 100 nm EPC-LRBPE (99:1)vesicles, or mixtures of 50 nm EPC-BCPE (8:2) vesicles were injectedinto the flow cell and incubated for various time periods (from 5 min to24 hr) in order to allow fusion and spreading on the sample surface toform a lipid bilayer. After incubation, the cell was rinsed with aphosphate buffer (0.05 M, pH 7.0) to remove unfused-lipid excess.

The smoothness and integrity of the supported bilayer formed on theMPS-APS-modified GaAs substrate were observed both byBioScope-AFM-surface-analysis measurement using 50 and 100 nmEPC-vesicles, and by fluorescence microscopy (Inverted microscope 1×70,Olympus, objective ×40) using 50 and 100 nm of fluorescently labeledEPC-LRBPE (99:1) mixture. AFM images of the lipid bilayers were acquiredin tapping mode in a phosphate buffer solution (0.05 M, pH 7.0), at roomtemperature (23-25° C.). The typical target amplitude was 300 mV (˜22nm) and set point was 230-240 mV (˜17 nm) for all measurements.Therefore, a “light” tapping was applied to avoid a possible damage tothe membranes and influence of the scanning itself on the membranestate. Image analysis performed using WSxM 5.0 Develop 1.2 software(Schoning and Poghossian, 2002).

Fluorescence imaging of the obtained bilayers indicates that in order toobtain a homogeneous surface coverage of the MPS-APS-modified GaAssubstrates, 50 nm vesicles should be used (FIG. 3, panels A, C). This isin contrast to the case of HF-etched glass slide where 100 nm vesiclesize is enough for uniform membrane formation (FIG. 3, panels B, D). Thesurfaces were illuminated with 532 nm light. It is well known thatunilamellar lipid vesicles can fuse and spread on surfaces such as glassto form homogeneous surface coverage of a lipid bilayer (Boukobza etal., 2001). Hence, we could compare the results on glass with those onMPS-APS-modified GaAs substrate. When the vesicles on the surface wereunruptured, the fluorescence image was non-homogeneous and looked grainy(FIG. 3, panel A). The image looks smooth when the sample is illuminatedwith a 532 nm beam in the absence of the vesicles. As further found,incubation of the EPC-LRBPE (99:1) vesicles on the MPS-APS-modified GaAssubstrate, for 5 min, was sufficient for vesicle adhesion and ruptureinto a bilayer membrane. However, some vesicles remained unruptured,leading to a grainy fluorescence image. Longer incubation time prior towashing with phosphate buffer increases the number of unrupturedvesicles. Downsizing the EPC-LRBPE (99:1) vesicles from 100 to 50 nmdramatically decreased the number of unruptured vesicles on theMPS-APS-coated GaAs substrate, leading to a homogeneous image similar tothat observed on an HF-etched glass slide (FIG. 3, panel C). Themembrane attachment to the MPS-coated GaAs devices in the presence ofAPS was improved compared to that in the absence of APS (data notshown).

The effect of vesicle size on the formation of a homogeneous bilayer onthe MPS-APS-modified GaAs MOCSER substrate was also evaluated usingAFM-surface-analysis measurement as shown in FIG. 4. The root meansquare roughness value of the ˜25 nm MPS-APS-modified GaAs devices wasfound to be ˜1.6 nm. The MPS polymer was not homogeneous and containedsome holes varying from 5 nm to 120 nm in diameter, sufficient to trap asingle 100 nm vesicle, preventing its attachment to nearby vesicles andtherefore its rupture owing to lack of certain surface density ofvesicles (Chai et al., 2002). By downsizing the vesicle size from 100 to50 nm, the possibility of two vesicles to adsorb to the same hole wasincreased, leading to their rupture into a homogeneous bilayer (Richteret al., 2006).

The stability and integrity of a membrane formed on MPS-APS-modifiedGaAs substrate over time were compared to those of a membrane formed ona glass, using fluorescence imaging. While integrity of the membraneadsorbed on HF-etched glass slide was stable for more than 7 days in thepresence of 2 mM CaCl₂, integrity of the membrane adsorbed on theMPS-APS-modified GaAs substrate deteriorated after ˜5 days (data notshown).

The results presented above show that GaAs device can be coated with aprotecting layer of MPS and that a uniform membrane can be formed on topof an MPS-APS-modified GaAs device.

Streptavidin Attachment to EPC-BCPE (8:2)-Based Membranes

Streptavidin attachment to EPC-BCPE (8:2) membranes was confirmed byBioScope-AFM-surface-analysis measurements as described above. For thispurpose, a flow cell was constructed on the surface of MPS-APS-modifiedGaAs samples by attaching HF-etched glass slide with double-sided tape.First, a lipid bilayer membrane was formed on MPS-APS-modified GaAssamples by incubating 50 nm EPC-BCPE (8:2) vesicles as described above.Next, 1 mg/ml streptavidin solution in phosphate buffer (0.05 M, pH 7.0)was added to the system, allowed to interact with the biotin moleculesfor 5 min so as to form biotin-streptavidin complex on top of thebiotinylated membrane, and was then rinsed with phosphate buffer. Insome experiments, 50 nm EPC-BCPE (498:1) vesicles (at a vesicleconcentration of 3×10⁸/μl) were added and allowed to interact for 5 minwith the surface-adsorbed streptavidin molecules, and were then rinsedwith phosphate buffer.

Anti-Hemoglobin Antibodies Attachment to MPS-APS-Modified MOCSER Surface

Since protein adsorption on confined surfaces is complicated, the GaAssurface was first modified by adsorption of MPS-APS layers as definedabove. Anti-human hemoglobin antibodies attachment to theMPS-APS-modified surface was then achieved by immobilizing saidantibodies on the surface through Protein G, and blocking thenon-binding sites by BSA, using spotting technique and fluorescent(fluorescein isothiocyanate, FITC)-labeled antibodies, as shown in FIG.5. A drop of 0.2 mg/ml of protein G in HEPES buffer (50 mM, pH 7.4) wasfirst placed on the surface ((a) in FIG. 5) and incubated for 10 minfollowed by a quick wash in a phosphate buffer (50 mM, pH 7.4). A biggerdrop of 0.1 mg/ml of BSA in HEPES buffer (50 mM pH 7.4) was placedcovering the area of the Protein G and a part of the bare GaAs surface((b) in FIG. 5), and was incubated for 5 min. After washing with aphosphate buffer (50 mM, pH 7.4), a solution of FITC-labeled anti-humanhemoglobin antibodies (0.1 mg/ml) in a phosphate buffer (50 mM, pH 7.4)was placed over the whole surface ((c,d) in FIG. 5), incubated for 10min and washed in a phosphate buffer (50 mM, pH 7.4). Care has beentaken that surface is not dried during the protein adsorption on thesurface. As shown in FIG. 5, the area covered with BSA and anti-humanhemoglobin antibodies only ((b) in FIG. 5) shows minimum fluorescence,indicating that BSA completely blocks the anti-human hemoglobinantibodies; the areas where bare GaAs or Protein G is present show aslightly higher fluorescence ((c,d) in FIG. 5); and the area coveredwith Protein G, BSA and anti-human hemoglobin antibodies ((a) in FIG. 5)shows the highest fluorescence, indicating that the antibodies areattached to the MPS-APS modified surface through Protein G, and thatblocking of the non-binding sites by BSA is possible.

FIG. 6 shows the change in current of the MOCSER after sequentiallyadsorbing Protein G, BSA and anti-human hemoglobin antibodies. As shown,the net change in current is negative when Protein G is introduced intothe sensing area; positive upon introducing of BSA; and negative duringanti-human hemoglobin antibodies interaction with the surface of thedevice. According to the theory capacitive sensing in ISFETs(Ghafar-Zadeh et al., 2010), when a negative charge is accumulated onthe surface of the device, it attracts positive charges in the gatedarea which in turn populates the conduction channel with more electronsleading to rise in current between source and drain, and vice versa. Incontrast, although Protein G, BSA, anti-hemoglobin antibodies as well ashemoglobin are all negatively charged protein molecules at pH 7.4, inthis experiment they showed a different behavior in terms of change incurrent, demonstrating that the sensing mechanism of the MOCSER isdifferent from most generally accepted capacitive theory applicable forISFETs.

Analytes and Solutions

3-mercaptopropyltrimethoxysilane (MPS; Cat. No. 63800) and3-aminopropyltrimethoxysilane (APS; Cat. No. 15629TU) were purchasedfrom Sigma. Egg phosphatidylcholin (EPC; Cat. No. 840051),1,2-dioleoyl-sn-glycero-3-phosphoethanolamine-N-(lissamine rhodamine Bsulfonyl) (LRBPE; Cat. No. 810150) and1,2-dioleoyl-sn-glycero-3-phosphoethanolamine-N-(cap biotinyl) (BCPE;Cat. No. 870273) were obtained from Avanti polar lipids, Alabaster,Ala., USA. L-lysine, L-glutamic acid, avidin, streptavidin, rabbitanti-streptavidin antibody, mouse anti-streptavidin antibody and variousphosphate buffer solutions with pH ranging from pH 6.0 to pH 8.0 wereused in this study. L-lysine (Cat. No. L5501), L-glutamic acid (Cat. No.G1251), streptavidin (Cat. No. S4762), IgG rabbit anti-streptavidinantibody (Cat. No. S6390), human hemoglobin, Streptococcus Protein G,and bovine serum albumin (BSA) were obtained from Sigma. Mousemonoclonal anti-streptavidin antibody (Cat. No. ab10020) was purchasedfrom Abcam. Sheep anti-human hemoglobin antibody (Cat. No. A80-135A) waspurchased from Bethyl Laboratories. Sodium phosphate monobasic (Cat. No.567545) and sodium phosphate dibasic (Cat. No. 567550) were obtainedfrom Merck KGaA, Darmstadt, Germany. Deionized Milli-Q water was usedfor buffer preparation and experiments. Urine used has been collectedonce from a person in the laboratory.

Experimental Procedure for the Study Described in Example 1

Experimental setup is shown in FIG. 7. The surface of the device wasmodified with MPS-APS layers as described above, and the chip containing16 MOCSERs was wire bonded for electrical measurements. All measurementswere performed on 4 MOCSERs simultaneously, using Keithley 236source-measure units and Keithley 2700 switch control, controlled andmonitored by Labview application (version 8.2).

Polydimethylsiloxane (PDMS)-based flow cell (4 mm in length and width,and 0.6 mm in height) was fixed on top of the sensing area with epoxyglue. The PDMS prepolymer was made from a mixture of RTV 615 siliconecompound and a curing agent (GE Silicones, Dandenong, Australia) at 10:1ratio. Transferring of analyte and buffer solutions to the MOCSERdevices was performed at 0.02 ml/min using a peristaltic pump (EP-1Econo pump, Bio-Rad Laboratories, Israel) with teflon pipes (innerdiameter of 0.8 mm).

Next, a lipid bilayer was formed on top of the device by introducing thelipids into the system with the peristaltic pump, as described above.After incubation, the sensing area was rinsed with a phosphate buffer.In order to preserve the membrane, the MOCSER devices were kept inliquid medium during the measurement. The analytes were dissolved in thephosphate buffer and were injected sequentially into the flow cell.Different concentrations were obtained by mixing the analytes withphosphate buffer solution (concentrations, volumes and flow-rates areprovided in Table 1). Phosphate buffer solution was injectedsequentially between analytes, using teflon pipes with an inner diameterof 0.8 mm, to rinse the sensing area and remove analyte excess. Thesignal measured during this time was used as a baseline in the dataanalysis. Phosphate buffer (0.05 M, pH 7.0) was used in all experimentsexcept in the case of monoclonal mouse anti-streptavidin antibody, wherephosphate buffer (0.05 M, pH 7.4) was used.

A constant potential of 1.0 V was applied between source and drain ofthe MOCSER devices, and the change in the source-drain current weremonitored as a function of time for all four selected channelssimultaneously. An Ag/AgCl pseudo reference electrode was placed in asealed tube and connected via a salt bridge to maintain a stable andconstant potential over the surface of the MOCSER devices.

TABLE 1 Analytes used in the study described in Example 1 AnalyteConcentration* Sample volume Flow rate Membrane L-lysine 1.6-50 μmole250 μl 0.02 ml/min 1) EPC 2) no membrane L-glutamic acid 0.4-50 μmole250 μl 0.02 ml/min 1) EPC 2) no membrane Phosphate buffer 0.05M 500 μl0.02 ml/min 1) EPC solutions with pH 2) no membrane ranging from 6.0 to8.0 Streptavidin/ 0.07-80 nmole 100 μl 0.01 ml/min 1) EPC-BCPE (8:2)Avidin 2) EPC 1) Rabbit 0.15-1 mg/ml 250 μl 0.01 ml/min 1) EPC-BCPE(8:2)- anti-streptavidin streptavidine antbody 2) EPC 2) Mouse 0.15-1mg/ml anti-streptavidin antbody 3) Rabbit serum — *Since a phosphatebuffer solution (0.05M, pH 7.0) was injected sequentially betweendifferent analytes to remove analyte excess, analytes were diluted withthe phosphate buffer. Therefore, amounts of analyte molecules and volumeare indicated separately.

Experimental Procedure for the Study Described in Example 2

Experimental setup is shown in FIG. 8. The surface of the device wasmodified with MPS-APS layers as described above, and the chip containing16 MOCSER devices was wire bonded for electrical measurements. Allmeasurements were performed on 4 MOCSERs simultaneously, using Keithley236 source-measure units and Keithley 2700 switch control, controlledand monitored by Labview application (version 8.2).

PDMS-based flow cell (4 mm in length and width, and 0.6 mm in height)was fixed on top of the sensing area with epoxy glue. The PDMSprepolymer was made from a mixture of RTV 615 silicone compound and acuring agent (GE Silicones, Dandenong, Australia) at 10:1 ratio.Transferring of analyte, buffer solutions and urine solutions to theMOCSER devices was performed at 0.02 ml/min using a syringe pump(Harvard Apparatus, PHD Ultra). Sensing of hemoglobin was performedthrough the binding of hemoglobin in the analyte introduced to itsantibodies.

A constant potential of 1.0 V was applied between source and drain ofthe MOCSER devices, and changes in source-drain current were monitoredas a function of time for all four selected channels simultaneously. AnAg/AgCl pseudo reference electrode was placed in a sealed tube andconnected via a salt bridge to maintain a stable and constant potentialover the surface of the MOCSER devices.

Example 1 The Response of the Membrane-Coated GaAs-Based MOCSER toVarious Analytes

FIG. 9 shows the change in the current through the MOCSER when thedevice is exposed to phosphate buffer solutions with pH ranging from 6.0to 8.0. As shown, following the time it takes the solution to reach thesensor, the MOCSER source-drain current response to pH change isimmediate and stable (FIG. 9A), and the response of the device to pH islinear within the range studied (FIG. 9B).

Changes in the source-drain current of the membrane-coated MOCSER devicewere observed when it was exposed to various concentrations ofnegatively- or positively-charged amino acids at pH=7, exemplified byL-glutamic acid (FIG. 10) or L-lysine (FIG. 11), respectively. As shown,the MOCSER source-drain current response is correlated with theconcentration of the analyte molecules. While the current increased asthe concentration of L-glutamic acid increased, it decreased withincreasing L-lysine concentration. The calibration plots are obtained byplotting the derivative of the signal as a function of time vs. theanalyte concentration (FIGS. 10B and 11B). The change in the signal uponexposure of the sensor to the analyte depends on the flow rate in themicrofluidic device; the specific response of the MOCSER to the analyte;and the analyte concentration. Since the flow rate was maintainedconstant through all the measurements and the only parameter varying wasthe analyte concentration, the derivative gradient of the signal shouldbe proportional to the analyte concentration. Indeed, the signalderivative as measured at about 200 sec after exposure to the analytesolution was found to be proportional to the concentration of theanalyte. Monitoring the gradient instead of the maximum current ensuresbetter reproducibility of the signal and eliminates contribution frombaseline shift.

The detection thresholds for L-lysine and L-glutamic acid in thepresence of EPC membrane were about 12.5 mM and 6.2 mM, respectively,and they improved to 3.2 mM and 1.6 mM for L-lysine and L-glutamic acid,respectively, in the absence of the membrane. These data further confirmthe existence of the membrane on the MOCSER. Clearly, the membranereduces the sensitivity of the device by about a factor of four.

It is interesting to note that while the change in the signal, in thecase of the pH measurements, is proportional to the logarithmic changein the concentration of the protons, it was almost linearly dependent onthe concentration of the amino acids. This phenomenon is not related tothe existence of the membrane and hence indicates that there is adifferent mechanism for the effect of both type of species (protons andorganic acids) on the MOCSER.

For streptavidin and rabbit anti-streptavidin antibodies detection, thestrategy presented in FIG. 12 was used, utilizing an EPC-BCPE (8:2)membrane containing a fraction of ˜20% biotin. Attachment ofstreptavidin to the membrane was evidenced and characterized by AFMusing 50 nm EPC-BCPE (498:1) vesicles as a marker. The 50 nm EPC-BCPE(498:1) vesicles were not observed in the absence of streptavidin.Exposing the membrane to either streptavidin or avidin at concentrationsabove 0.8 μM at pH=7 resulted in a significant change in the MOCSERsource-drain current. The current increased upon exposure tostreptavidin concentration increase (FIG. 13A), but the signal did notrecover when the solution was changed to buffer with no streptavidin,indicating a strong (and seemingly irreversible on the time scale of theexperiment) binding of the streptavidin to the biotin. When exposed toavidin, the current through the MOCSER was reduced (FIG. 13B). When amembrane without biotin (EPC-based membrane) was exposed to the samesolution, a change in the current was observed; however, this changecould be completely reversed by washing with buffer.

While streptavidin is negatively charged at neutral pH, avidin ispositively charged. Hence a reverse effect on the current is observed inaccordance with the observations when amino acids were probed, generallyindicating that a negatively charged analyte causes an increase in thecurrent through the MOCSER while a positively charged analyte causes adecrease in the current.

Change in the MOCSER source-drain current was observed when devices towhich streptavidin were initially attached were exposed to polyclonalrabbit anti-streptavidin antibody in serum (FIG. 14A, right). In thiscase, the current decreased upon exposure to the antibodies atconcentrations of 0.031, 0.125 and 1 mg/ml. Irreversible negative offsetin the current was observed following washing with a buffer solutionindicating binding of the anti-streptavidin molecules to thebiotin-streptavidin complex. However, when the serum contained noantibodies, the signal returned to the baseline upon washing (FIG. 14A,left). The fact that the signal is not returning to the baseline whenthe antibodies are present reflects the strong binding of the antibodyto the streptavidin. As a control experiment, devices with no biotin(EPC-based membrane) and hence no bound streptavidin were exposed to thesame antibody containing-serum solution. A small positive offset in thecurrent was observed upon washing, indicating non-specific binding ofserum species to the membrane. The normalized response of the current isplotted in FIG. 14C as a function of the amount of antibodies to whichthey were exposed. The amount of analyte (and not its concentration) ispresented since the binding of the analyte is seemingly irreversible andthe signal is accumulating as a function of the amount of analyte towhich the sensor is exposed. The two curves relate to devices coatedwith biotinylated membrane (EPC-BCPE (8:2) based membrane) to whichstreptavidin was attached and to devices coated with a membrane(EPC-based membrane) without biotin.

For comparison, FIG. 14B shows the normalized change in the currentobtained when the device was exposed to various concentrations ofmonoclonal mouse anti-streptavidin antibody. In this case, the signalincreased upon exposure to the antibodies and exhibited a baselineoffset upon washing with the buffer solution. The difference in thetrend of the response (negative vs. positive offset in FIGS. 14A and14B, respectively) is probably due to different charge on each type ofantibody. The normalized response of the current through the MOCSER dueto interaction of mouse anti-streptavidin antibody with streptavidin isshown in FIG. 14D. Here, clearly a nonlinear response of the device tothe concentration is observed. The non-linearity in the response is notevident in other cases (FIG. 14C) because of the limited range of theanalyte concentration to which the sensor was exposed and due to the lownumber of data points that do not allow a nonlinear fit.

The results presented above indicate that the operation of the MOCSER asa sensor is based on the fact that it is capacitance sensitive. Thus,when the device is immersed in electrolyte solution with a referenceelectrode, a double layer is formed on its surface, as shown in FIG. 15.Clearly, when the analyte on the surface of the membrane is negativelycharged, the charge accumulating on the surface of the device ispositive and vice versa. Since the device is based on n-doped GaAs,positive charge on the surface increases the charge carrierconcentration in the conductive channel and the source-drain currentincreases. The opposite is true for negative charge on the surface ofthe GaAs that causes depletion in the charge carrier concentration andhence reduction in the source-drain current.

Example 2 The Membrane-Coated GaAs-Based MOCSER is Capable of Detectingthe Presence of Hemoglobin in Urine

In this study, phosphate buffer (50 mM, pH 7.4) was first used as amodel system, i.e., as a carrier buffer for the hemoglobin, and wasinjected sequentially between analytes to rinse the sensing area andremove analyte excess. Then, hemoglobin was dissolved in urine, andurine was injected between analytes to rinse the sensing area. FIG. 16shows the change in the source-drain current through the MOCSER when thedevice was exposed to phosphate buffer-based hemoglobin solutions withconcentrations of 0.1, 0.5, 1, 5, 10 and 25 mg/ml, and FIG. 17 shows thechange in the source-drain current through the MOCSER when the devicewas exposed to urine-based hemoglobin solutions with concentrations of0.25, 0.5 and 1 mg/ml. As shown in FIGS. 16 and 17, the device responseis immediate and stable upon exposure to the hemoglobin solutions. Thecurrent measured in the MOCSER decreases when hemoglobin interacts withthe anti-hemoglobin antibodies attached to the MPS-APS-modified surface,and recovers after washing hemoglobin with the phosphate buffer orurine, wherein the signal is correlated with the concentration of theanalyte molecules. A calibration plot, obtained by plotting the slope ofthe signal as a function of time vs. the analyte concentration, is shownin FIG. 18.

Since all other parameters during the measurement were kept constant andonly the concentration of analyte was changed, the slope of the signalshould be proportional to the concentration of the analyte. Monitoringthe gradient instead of the net change in the current ensures betterreproducibility of the signal and eliminates contribution from baselineshifts. As found, the sensitivity of the MOCSER to hemoglobin, based onthese experiments, was 10 μg/ml and 100 μg/ml of hemoglobin in phosphatebuffer and urine, respectively, representing the lower detection limitfor the current setup. The sensitivity of the MOCSER to hemoglobin inurine was lower than in phosphate buffer, probably due to the fact thatthe urine salt concentration is much higher.

The high selectivity of the MOCSER results from the antibody-antigeninteraction. Thus, in order to verify the selectivity of the MOCSER, thesource-drain current in response to hemoglobin solutions in phosphatebuffer, when sheep anti-human hemoglobin antibodies are not immobilizedon the gate area, and the surface is only functionalized with protein Gand BSA, was measured. As found and shown in FIG. 19, no response to thehemoglobin analytes was observed, indicating the high selectivity of theMOCSER, and these results were consistent when current was measured inresponse to urine-based hemoglobin solutions. Nevertheless, when a veryhigh concentration of hemoglobin was used, a non-specific response wasobserved, but that response was completely random and could not becorrelated with the hemoglobin concentration. In order to verify thespecificity of the MOCSER, analytes containing avidin, representing anon-specific antigen, were introduced to a MOCSER having a gate area onwhich sheep anti-human hemoglobin antibodies were immobilized, and asshown in FIG. 20, no change in the current was observed, demonstratingthe high specificity of the device.

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1. A semiconductor device for the detection of an active site-containingprotein or a ligand thereof in a solution, said device being composed ofat least one insulating or semi-insulating layer, at least oneconducting semiconductor layer, two conducting pads, a protectivemolecular layer, and said ligand or active site-containing protein,wherein said at least one conducting semiconductor layer is on top ofone of said insulating or semi-insulating layers, said two conductingpads are on both sides on top of an upper layer which is either one ofsaid conducting semiconductor layers or another of said insulating orsemi-insulating layers, making electrical contact with said at least oneconducting semiconductor layer, said protective molecular layer isfabricated on top of said upper layer protecting said upper layer fromcorrosion, and said ligand or active site-containing protein is linkedeither directly or indirectly to said protective molecular layer,wherein exposure of said ligand or active site-containing protein, to asolution containing said active site-containing protein or ligand,respectively, causes a current change through the semiconductor devicewhen a constant electric potential is applied between the two conductingpads.
 2. A semiconductor device according to claim 1, composed of atleast one insulating or semi-insulating layer, one conductingsemiconductor layer, two conducting pads, a protective molecular layer,and said ligand or active site-containing protein, wherein saidconducting semiconductor layer is on top of one of said insulating orsemi-insulating layers, said two conducting pads are on both sides ontop of an upper layer which is either said conducting semiconductorlayer or another of said insulating or semi-insulating layers, makingelectrical contact with said conducting semiconductor layer, saidprotective molecular layer is fabricated on top of said upper layer, andsaid ligand or active site-containing protein is linked either directlyor indirectly to said protective molecular layer.
 3. A semiconductordevice according to claim 1, for the detection of said activesite-containing protein, wherein said device comprises said ligandlinked either directly or indirectly to said protective molecular layer,and exposure of said ligand to a solution containing said activesite-containing protein causes a current change through thesemiconductor device when a constant electric potential is appliedbetween the two conducting pads.
 4. A semiconductor device according toclaim 1, for the detection of said ligand, wherein said device comprisessaid active site-containing protein linked either directly or indirectlyto said protective molecular layer, and exposure of said activesite-containing protein to a solution containing said ligand causes acurrent change through the semiconductor device when a constant electricpotential is applied between the two conducting pads.
 5. A semiconductordevice according to claim 1, wherein (i) each one of said at least oneconducting semiconductor layer independently is a semiconductor selectedfrom the group consisting of a III-V and a II-VI material, or and amixture thereof, wherein III, V, II and VI denote the Periodic Tableelements II=Ga, In; V=As, P; II=Cd, Zn; VI=S, Se, Te. (ii) each one ofsaid at least one insulating or semi-insulating layers independently isa dielectric material selected from the group consisting of siliconoxide, silicon nitride and an undoped semiconductor selected from thegroup consisting of a III-V and a II-VI material, and a mixture thereof,wherein III, V, II and VI denote the Periodic Table elements III=Ga, In;V=As, P; II=Cd, Zn; VI=S, Se, Te.
 6. A semiconductor device according toclaim 5, wherein each one of said at least one conducting semiconductorlayer is doped GaAs or doped (Al,Ga)As; or said undoped semiconductor isundoped GaAs or undoped (Al,Ga)As. 7-8. (canceled)
 9. A semiconductordevice according to claim 1, wherein said protective molecular layercomprises an alkoxysilane-based polymer formed by polymerization ofdialkoxysilanes, trialkoxysilanes or tetraalkoxysilanes, each having afunctional group, a biotinylated form thereof, or a mixture of theaforesaid.
 10. A semiconductor device according to claim 9, wherein saidpolymer is formed by polymerization of dialkoxysilanes ortrialkoxysilanes of the general formula (C₁-C₇ alkyl)₂-Si(OR)₂ or (C₁-C₇alkyl)-Si(OR)₃, respectively, biotinylated forms thereof, or mixtures ofthe aforesaid, wherein each of the Rs independently is a (C₁-C₄)alkyl,preferably methyl or ethyl, and the (C₁-C₇)alkyl group of thetrialkoxysilane, or one or two of the (C₁-C₇)alkyl groups of thedialkoxysilane, is substituted at a terminal carbon atom with afunctional group selected from the group consisting of mercapto, amino,and hydroxyl; and is optionally further interrupted with one or more—NH— groups.
 11. A semiconductor device according to claim 10, whereinsaid polymer is formed by polymerization of (i) a mercapto-functionalalkoxysilane of the general formula HS—(C₁-C₇)alkylene-SiR(OR)₂ orHS—(C₁-C₇)alkylene-Si(OR)₃, preferably HS—(C₁-C₇)alkylene-Si(OR)₃, abiotinylated form thereof, or a mixture of the aforesaid, wherein eachof the Rs independently is a (C₁-C₄)alkyl, preferably methyl or ethyl;(ii) an amino-functional alkoxysilane of the general formulaH₂N—(C₁-C₇)alkylene-SiR(OR)₂ or H₂N—(C₁-C₇)alkylene-Si(OR)₃, preferablyH₂N—(C₁-C₇)alkylene-Si(OR)₃, a biotinylated from thereof, or a mixtureof the aforesaid, wherein each of the Rs independently is a(C₁-C₄)alkyl, preferably methyl or ethyl, and the C₁-C₇ alkylene isoptionally interrupted with one or more —NH— groups; or (iii) a mixtureof said mercapto-functional alkoxysilane and said amino-functionalalkoxysilane.
 12. A semiconductor device according to claim 11, whereinsaid mercapto-functional alkoxysilane ismercaptomethylmethyldiethoxysilane, mercaptomethylmethyldimethoxysilane, 3-mercaptopropylmethyldiethoxysilane,3-mercaptopropyl methyldimethoxysilane, 3-mercaptopropyltrimethoxysilane(MPS), or 3-mercapto propyltriethoxysilane; and said amino-functionalalkoxysilane is N¹-(3-(trimethoxysilyl)propyl)ethane-1,2-diamine,N′-(3-(triethoxysilyl)propyl)ethane-1,2-diamine,3-aminopropyltrimethoxysilane (APS),3-aminopropyltriethoxysilane,4-aminobutyltriethoxysilane, 4-aminobutyltrimethoxysilane,N¹-(3-(dimethoxy(methyl)silyl)-2-methylpropyl)ethane-1,2-diamine,N¹-(3-(diethoxy(methyl)silyl)-2-methylpropyl)ethane-1,2-diamine,aminopropylmethyldimethoxysilane, or aminopropylmethyldiethoxysilane.13. A semiconductor device according to claim 12, wherein said polymeris formed by polymerization of a mixture of MPS and APS.
 14. Asemiconductor device according to claim 1, wherein said ligand or activesite-containing protein is indirectly linked to said protectivemolecular layer via (i) a mono- or bi-layer membrane comprising anamphiphilic compound or a mixture thereof, wherein said mono- orbi-layer membrane is adhered to said protective molecular layer; or (ii)a linker selected from the group consisting of a ligand-binding proteinsuch as Protein A, Protein G, streptavidin, avidin or an antibody,biotin, and a biotin-like molecule.
 15. A device according to claim 14,wherein said ligand or active site-containing protein is (i) adsorbed toor incorporated into said mono- or bi-layer membrane; or (ii) linked tosaid protective molecular layer via a ligand-binding protein. 16.(canceled)
 17. A semiconductor device according to claim 14, whereinsaid amphiphilic compound is a phospholipid selected from the groupconsisting of a phosphoglyceride or phosphosphingolipid, a biotinylatedform thereof, and a mixture of the aforesaid.
 18. A semiconductor deviceaccording to claim 17, wherein said phosphoglyceride is selected fromthe group consisting of a plasmalogen, a phosphatidate, aphosphatidylethanolamine, a phosphatidylcholine such as eggphosphatidylcholin (EPC), a phosphatidylserine, a phosphatidylinositol,phosphatidylinositol phosphate, phosphatidylinositol bisphosphate,phosphatidylinositol triphosphate, a glycolipid such as aglyceroglycolipid, glycosphingolipid, and glycosylphosphatidylinopsitol,a phosphatidyl sugar, and a biotinylated form thereof such asdioleoyl-sn-glycero-3-phosphoethanolamine-N-(cap biotinyl) (BCPE),Biotin-Phosphatidylcholine, Biotin Phosphatidylinositol 3-phosphate,Biotin Phosphatidylinositol 4,5-bisphosphate, Biotinylatedphosphatidylinositol 3,4,5-trisphosphate, and1-((1-octanoyl-N′-biotinoyl-1,6-diaminohexane-2R-octanoyl)phosphatidyl)inositol-3,4,5-triphosphate,tetrasodium salt; and said phosphosphingolipid is selected from thegroup consisting of a ceramide phosphorylcholine, a ceramidephosphorylethanolamine, a ceramide phosphorylglycerol, and abiotinylated form thereof such as Biotin Sphingomyelin.
 19. Asemiconductor device according to claim 1, wherein said solution is anaqueous solution, a physiological solution, a bodily fluid such asamniotic fluid, aqueous humour, vitreous humour, bile, blood serum,breast milk, cerebrospinal fluid, cerumen (earwax), endolymph,perilymph, female ejaculate, gastric juice, mucus, peritoneal fluid,saliva, sebum (skin oil), semen, sweat, tears, vaginal secretion, vomitand urine, or a bodily fluid-based solution.
 20. (canceled)
 21. A methodfor detection of an active site-containing protein or a ligand thereofin a solution, said method comprising: (i) exposing a semiconductordevice according to claim 1 to said solution; and (ii) monitoring thepresence of said active site-containing protein or ligand in saidsolution according to the changes in the current measured in saidsemiconductor device when a constant electric potential is appliedbetween the two conducting pads.
 22. The method of claim 21, forquantification of said active site-containing protein or ligand thereofin said solution, wherein the current change is proportional to theconcentration of said active site-containing protein or ligand thereofin said solution.
 23. The method of claim 21, for studyingreceptor-ligand pair interactions, in particular, monitoring theinteraction of a receptor in a solution with a ligand linked eitherdirectly or indirectly to said protective molecular layer, or viceversa.
 24. The method of claim 21, wherein (i) said activesite-containing protein is an antibody, and said ligand is an antigen,or vice versa; (ii) said active site-containing protein is an enzyme,and said ligand is a substrate or inhibitor, or vice versa; (iii) saidactive site-containing protein is a receptor, and said ligand is aprotein or organic molecule, or vice versa; or (iv) said activesite-containing protein is a lectin, and said ligand is a sugar.